This invention relates to a cerium activated phosphor for computerized tomography and other x-ray, gamma radiation, and nuclear radiation detecting applications. More specifically, the invention relates to a lanthanide oxide fluoride compound having cerium luminescence.
Computerized tomography scanners are diagnostic instruments used in industrial and medical imaging. A body is exposed to a relatively planar beam or beams of radiation, such as x-ray or gamma ray radiation, the intensity of which varies in direct relationship to the energy absorption along a plurality of body paths. By measuring the radiation intensity along these paths from a plurality of different angles or views, a radiation absorption coefficient can be computed for various areas in any plane of the body through which the radiation passes. The absorption coefficients are used to produce a display of, for example, bodily organs intersected by the radiation.
Phosphors can be used to form scintillators which are excited by the impinging X-ray or gamma radiation, and emit optical wave length radiation. The optical output from the scintillator material is made to impinge upon photo electrically responsive materials in order to produce electrical output signals. The amplitude of the signals is directly related to the intensity of the impinging X-ray or gamma radiation. The electrical signals are digitized for processing by digital computer means which generate the absorption coefficients in a form suitable for display on a cathode ray tube screen or other permanent media.
In general, it is desirable that the amount of light output from the phosphors and resulting scintillator be as large as possible for a given amount of X-ray or gamma ray energy. This is particularly true in the medical tomography area where it is desired that the energy intensity of the X-ray be as small as possible to minimize any danger to the patient.
Another important property that the phosphor material should possess is a short primary decay, which can be measured as a decay time constant. As used herein, the term "decay time constant" means the time for luminescence output to decay to about 36.8 percent of the maximum light output after the excitation radiation ceases. This means that there should be a relatively short period of time between the termination of the high energy radiation excitation and the cessation of light output from the phosphor or scintillator. If this is not the case, there is a blurring, in time, of the information bearing signal generated, for example, when the scintillator is used to produce tomographic imaging data. Furthermore, if rapid tomographic scanning is desired, the presence of the primary decay tends to severely limit the scan rate, thereby rendering difficult the view of moving bodily organs, such as the heart or lungs.
Positron emission tomography scanners utilize gamma ray detector systems. The detector system is capable of capturing gamma rays and converting them into a luminescent output. The luminescent output is converted by means of a photo multiplier into an electrical signal. Bismuth germanate has the necessary high stopping power required for capturing gamma radiation, and has been used in gamma ray detection systems. The gamma ray stopping power of bismuth germanate has been measured to be about 0.955 per centimeter. The decay time constant for bismuth germanate is about 300 nanoseconds.
It is an object of this invention to provide a phosphor comprised of cerium in a lanthanide oxide fluoride compound wherein the lanthanides are lutetium, gadolinium, yttrium, and mixtures thereof having a tetragonal crystal form.
It is another object of this invention to provide a phosphor having a high X-ray or gamma ray stopping power, e.g., comparable to bismuth germanate.
It is another object of this invention to provide a phosphor having a high X-ray or gamma ray stopping power, and a fast decay time constant, e.g., less than the 300 nanosecond decay time constant for bismuth germanate.